In sonographic systems, acoustic signals are transmitted by a scanhead into a body or other subject and reflected signals are received by the scanhead for image processing. The reflected signals are used by the sonographic system to form images of the structure of the body matter (e.g., a patient's tissue) or other subject of interest. A scanhead used in such sonographic imaging is typically a hand held enclosure which contains one or more independent transducers and possibly other electronics.
The transducers of a sonographic system scanhead convert electrical energy to mechanical (acoustic) energy radiating away from its surface when transmitting and mechanical (acoustic) energy impinging upon its surface to electrical energy when receiving. An individual portion of transduction material is called an element which is often manufactured as a particular geometric shape such as a rectangle. Typically, these transducer elements are arranged in a regular pattern (an array) with their centers arranged as a line to form a linear array or phased array, along an arc to form a curved array, or in a grid to form a 2D array. Usually this regular pattern of transducer elements has a repeated spacing as measured from element center to element center which is called the pitch. In sonographic imaging operations transducer elements are generally used in groups. The total extant of such a group of transducer elements in a dimension is the aperture in that dimension. For example, for a linear array, one dimension is the height of the transducer element while the other dimension is the number of transducer elements used times the pitch.
An ultrasonic beam may be formed, whether in transmit or receive operation, through appropriate use of the foregoing groups of transducer elements. For example, a receive beam is formed by adjusting one or more attributes of the transducer element signals (e.g., delaying and/or weighting to provide transducer element beamforming signals corresponding to transducer elements of the selected aperture) and summing these transducer element beamforming signals to provide a beamformed signal having a maximum signal response corresponding to a particular point (the particular point being a “focal point”). The foregoing transducer element signal attributes are referred to herein as beamforming parameters. Such beamforming parameters are typically utilized to form beams to reject clutter (e.g., undesired reflected signals etc.) received from undesired areas (e.g., directions other than a desired “look direction”).
In particular, the delays are applied to the transducer element signals from the group of transducer elements such that if a narrow pulse were emitted from the focal point, the signals having been thus delayed would arrive at a summing device at the same time and therefore would result in the largest value. This same narrow pulse coming from any other point than the focal point would not arrive simultaneously at the summer and therefore would not sum to be as large a signal. A beam having a particular shape (e.g., width, length, direction, etc.) may be formed through use of appropriate beamforming parameters. For example, a mainlobe that is “pointed” in a desired “look” direction may be formed.
Independent from the application of delays to create beams, an aperture may be apodized. Apodization is the process of applying potentially unique gain values (weighting) to the transducer element signals before they are summed. Particular apodization functions may be applied to apertures for generating beams having desired attributes, such as reduced sidelobes and thus to further reject clutter. There are many standard weighting functions that can be applied to an aperture, but there are three that are particularly exemplary. These are uniform weighting (also known as rectangular, box car, sinc, or unapodized), Hanning (also known as Hann) weighting, and cosine weighting. Hanning weighting (1+cos(x)) and cosine weighting (cos(x)) are related to one another in that Hanning weighting is a raised cosine function. Other mathematical functions which may be used for aperture apodization in ultrasound imaging systems are Hamming, Blackman-Harris, or other application specific window functions.
The beam formed by uniformly weighted aperture is termed a Sinc beam, the beam formed by Hanning weighted aperture is termed Hanning beam, and the beam formed by cosine function weighted aperture is termed cosine apodized beam. An object is scanned by sequentially shifting an ultrasound beam (e.g., Sinc beam, Hanning beam, or Cosine apodized beam) to form an image. Depending upon the implementation, an ultrasound image can be formed either by Sine beam or Hanning beam or beams of other types.
When beamforming parameters (e.g., delays) are continuously adjusted so that the focal point moves along a particular direction, a dynamically focused beam is created. In providing beam scanning for sonographic imaging, these dynamic beams are usually formed so that the focal point follows a straight line in Cartesian space for linear arrays or along a single angle from an apex in either phased or curved arrays. For example, by sequentially adjusting the beamforming parameters of the transducer element signals a series of beams may be formed to scan a volume of interest (e.g., a particular area or depth within a patient may be scanned). Information from a plurality of such scanned beams can be aggregated to generate an image of the scanned volume of interest (e.g., an ultrasound image of a sub-dermal portion of a patient). For example, in ultrasound B-mode operation, an image is generated from multiple lines of echo data received from a plurality of ultrasound beams of different look directions (e.g., beams scanned in different look directions). Such image generation from scanned beams is referred to herein as scanned volume imaging.
It is known that the monochromatic signal acquired by the Hanning beam is mathematically equal to summing the signal acquired from the Sinc beam with the average of signals acquired from two spatially shifted neighboring Sinc beams, provided that these beams are spaced according to Nyquist theorem. That is, the first null of the left Sinc beam and the first null of the right Sinc beam must be aligned with the peak of the center Sinc beam. Based on these properties, by processing signals acquired from three adjacent Nyquist spaced Sinc beams, a technique has been proposed to improve the performance in radar applications. However, in ultrasound imaging, the line density is selected according to multiple system parameters for optimal image quality, thus setting the sampling spacing from beam to beam or scan line to scan line according to Nyquist criterion generally cannot be satisfied. Furthermore, in ultrasound imaging dynamic beamforming, as may be used in providing the aforementioned scanned sample beams, is usually implemented in conjunction with a variable aperture. In other words, different aperture sizes are used for forming beams at different depths. Thus, the clutter reduction technique based on processing Sinc beam, such as may be implemented for radar, often cannot be adopted for use in ultrasound scanned volume imaging.
FIG. 1A illustrates the aforementioned scanned volume imaging. Specifically, transducer 11, having transducer elements E1 to EN, shown in FIG. 1A may be operated to provide such scanned volume imaging. In operation, transducer element signals of transducer elements E1 to EN are processed to form receive beams directed to particular areas within volume being imaged 15. Such beams may be formed to collect information regarding objects (also referred to as objects of interest) within volume being imaged 15, such as object 12 (e.g., fluid filled region) and object 13 (e.g., tissue structure) present below surface 16 (e.g., skin surface).
It should be appreciated that the higher the signal to clutter ratio in the signals (e.g., beamformed signals) used in scanned volume imaging, the higher the contrast resolution (e.g., better tissue differentiation) will be in the generated image. One source of signal clutter are the aforementioned sidelobes which typically accompany the mainlobes of the generated beams. The presence of undesired sidelobes in association with desired mainlobes can be seen from the illustration of FIG. 1A. Specifically, the mainlobes illustrated in FIG. 1A each have sidelobes associated therewith (e.g., sidelobes SL5 associated with mainlobe ML5, the combination of which are shown in a dashed line portion to help in distinguishing these lobes from the composite representation). The number and level of the sidelobes and their structure define how much of the off-axis undesired echoes are integrated into the resulting beamformed signal, thus cluttering the desired echoes of the object of interest. The ability to reduce the sidelobes improves the contrast resolution or the differentiability of objects of interest, such as tissues in an image.
Another source of image degradation is the width of the mainlobes used to collect image information. For example, the width of the mainlobe defines how an object within a volume being imaged is spread by the beam. Thus, the width of the mainlobe typically relates to the detailed resolution of an image. Accordingly, it is often desirable that the beams formed for the aforementioned scanning have a narrow focus so that objects of interest in the generated images can be well defined.
From the above it can be appreciated that the width of the mainlobe, the level of the sidelobes, and the structure of the sidelobes (e.g., how fast the sidelobes roll off from the mainlobe) have great significance to image quality. For example, higher resolution images can be achieved with very well-defined beams.
Signal processing for image generation using transducer 11 of FIG. 1A may include forming beams using a selected aperture (e.g., a selected group of transducer elements, such as transducer elements E06-E15) by appropriately implementing beamforming parameters (e.g., delays and/or weights) for the transducer element signals received by the transducer elements of the selected aperture. For example, delays of the beamforming parameters may be selected to provide mainlobes ML11 having desired focal points (e.g., applying appropriate delays to provide beams to scan a particular depth of volume being imaged 15). Additionally, the beamforming process may involve applying appropriate weights (apodization process) to the signals received from the transducer elements of the selected aperture, such as to reduce sidelobes associated with the mainlobes. Thus, the beam forming parameters utilized in generating beams may comprise complex values such that the signal received from transducer elements may be modified both in magnitude and phase.
Although generally reducing the sidelobes of the beam, use of the aperture apodization process spreads the mainlobe. Undesirable results associated with the use of the foregoing typical beamforming using an apodization processes are illustrated by FIGS. 1B-1D. FIG. 1B shows tissue mimic phantom 150 (generally representing volume being imaged 15 shown in FIG. 1A) which is composed of fluid filled region A on the left (such as may correspond to a portion of the volume being imaged of FIG. 1A comprising object 12) and tissue region B on the right (such as may correspond to a portion of the volume being imaged of FIG. 1A comprising object 13). Tissue region B is assumed to comprise a cluster of point scatterers (e.g., point scatters 14) of equal scattering cross-sections. An image is formed when a volume being imaged, represented here by tissue mimic phantom 150, is insonified with a sequence of ultrasound beams formed by the linear array of elements E1-EN. Since little scattering intensity will be received from fluid filled region A, the resulting image (in an ideal situation) would hold no gray scale displayed for fluid filled region A, whereas tissue region B would display a distribution of dots with similar intensity to those shown in the mimic phantom.
As discussed above, in conventional ultrasound imaging systems, the aperture of an array is either apodized with a deterministic mathematical function to partially suppress the sidelobes (thereby widening the mainlobe) for improvement of image contrast or not apodized to maintain a narrower mainlobe thereby yielding smaller imaging dot size with increased clutter. Each results in a degradation of image quality and will display a distorted image.
FIG. 1C shows the two different beam configurations discussed above to illustrate the problem. Beam BU is an unapodized beam (e.g., a Sinc beam formed using a uniform weighting function to define beamforming weighting distribution) providing a more narrow mainlobe having sidelobes with relatively high levels. Beam BH is an apodized beam (e.g., a Hanning beam formed using a raised-cosine weighting function to define beamforming weighting distribution) providing a more wide (spread) mainlobe having sidelobes with relatively low levels. The magnitude of the beams illustrated in FIG. 1C are logarithmically compressed and the sidelobes are scalloped and gradually roll off.
It is assumed that beams BU and BH are used to image a same area, specifically a portion of tissue region B of tissue mimic phantom 150 of FIG. 1B. Beam BU generates object of interest representation 101 (as may be used in aggregating an image generated by scanning a plurality of beams BU in different look directions within the area represented by tissue mimic phantom 150) resulting from reflected signals (e.g., reflected by point scatters 14) received by the mainlobe. Beam BU further generates artifacts 101-1 to 101-8 (also as may be aggregated into a generated image as undesired clutter) resulting from reflected signals received by the sidelobes. Likewise, beam BH generates object of interest representation 100 (as may be used in aggregating an image generated by scanning a plurality of beams BH in different look directions within the area represented by tissue mimic phantom 150) resulting from reflected signals (e.g., reflected by point scatters 14) received by the mainlobe. Beam BH further generates artifacts 100-1 to 100-4 (also as may be aggregated into a generated image as undesired clutter) resulting from reflected signals received by the sidelobes. As can be seen in FIG. 1C, although the same area of an object was imaged, object of interest representation 100 as provided by beam BH is spread compared to object of interest representation 101 provided by beam BU. Also as can be seen in FIG. 1C, more (albeit smaller) artifacts are generated by beam BU (artifacts 101-1 to 101-8) than artifacts generated by beam BH (artifacts 100-1 to 100-4).
A sonographic image may be generated by scanning a plurality of either beam BU or beam BH to insonify a volume being imaged. For example, the representations created by scanning a respective one of beams BU and BH throughout the area represented by tissue mimic phantom 150 may be aggregated to form an image of an object of interested. However, as can be appreciated from the illustration of FIG. 1C, when using beam BU the object of interest in the generated image may be relatively sharp because the object of interest representations (e.g., object of interest representation 101) are relatively small but the number of artifacts (e.g., artifacts 101-1 through 101-8) is high as a result of the more prominent sidelobes. The artifacts associated with the use of beam BU also extend a long distance from corresponding ones of object of interest representations, further degrading the generated image. Also as can be appreciated from the illustration of FIG. 1C, when using beam BH the object of interest in the generated image is less sharp because the object of interest representations (e.g., object of interest representation 100) are relatively large but the number of associated artifacts (e.g., artifact 100-1 through 100-n) is low as a result of the less prominent sidelobes. Moreover, the artifacts associated with the use of beam BH extend a shorter distance from the object of interest representation. Each of the foregoing beam forming techniques, therefore, results in generated images which often are of a lower quality than desired. As can be appreciated from the foregoing, achieving a well-defined beam without significant sidelobes for providing quality imaging has proven illusive.